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BACKGROUND OF THE INVENTION
Ventricular fibrillation (VF) is a lethal cardiac arrhythmia for which the
only known efficacious treatment is electrical countershock. A victim of
VF outside of the hospital setting has little chance of survival since
treatment must take place within a few minutes after the onset of the
episode.
Fortunately, new techniques and devices are being devised to help deal with
this life threatening condition. Among these are computer techniques which
aid in the identification of high risk VF patients, anti-arrhythmic drugs
which can prophylactically be administered to these patients, programs for
wide-spread cardiopulmonary resuscitation training, and implantable
devices which can automatically detect VF and deliver cardioverting
countershocks.
"Cardioverting" or "cardioversion" as used herein is intended to encompass
the correction of a number of arrhythmic heart conditions, both lethal and
non-lethal. Those arrhythmic heart conditions include atrial tachycardia,
atrial flutter, atrial fibrillation, junctional rhythms, ventricular
tachycardia, ventricular flutter, ventricular fibrillation, and any other
non-pacemaking related arrhythmic condition which may be corrected by
applying electrical shocks to the heart. Obviously then, "defibrillation"
is included in the term cardioversion as a method of applying electrical
shocks to the heart to defibrillate fibrillating atria or fibrillating
ventricles.
Many of the known techniques, such as defibrillation in a hospital setting
or defibrillation by a paramedic as part of a resuscitation program, rely
upon the human detection of VF. This detection has typically been
accomplished by a trained operator interpreting an ECG from an
oscilloscope tracing. However, there are situations where such an approach
to reversing VF is impossible or impractical. There is accordingly a great
need for an electronic device able to accurately detect VF or other life
threatening arrhythmias from an input ECG where such a traditional
approach is unfeasible. For example, an external defibrillator could be
built with an interlock to its discharge switch so that a shock can be
delivered only after the presence of VF has been confirmed by a detector
receiving an ECG signal from the paddles. Such a defibrillator could
safely be used by even an untrained operator.
With regard to the automatic implantable defibrillator, techniques have
been developed which are generally acceptable for detecting VF and
discriminating between life threatening arrhythmias and other cardiac
malfunctions. Yet there is considerable room for improvement with regard
to detecting and discriminating VF from other non-fatal arrhythmias.
Accordingly, another use for such a detector as noted above would be in
the fully-implantable automatic defibrillator.
Previous approaches to VF detection for implantable devices have had
certain drawbacks. Fundamental questions, particularly important to an
automatic implantable defibrillator, relate to potential failure modes,
the risks to a patient should the device reach one of these failure modes,
and specifically to whether failure should occur in a passive or an active
manner. Considerations of failure modes in another field, for example,
have led pacer manufacturers to design pacers to resort to fixed rate
pacing, an active mode, should there be a sensing failure such as caused
by interference. The risk of competing with the natural heartbeat has been
judged less than the risk of potential inhibition when pacing is needed.
Similar principles apply to the automatic implantable defibrillator, though
simple answers do not exist. An active mode failure would result in the
delivery of a shock when none is necessary, an occurrence which could be
particularly unpleasent to the patient. A passive mode failure, on the
other hand, would inhibit the delivery of a necessary shock, and could
result in death. Obviously, failures must be minimized, but they still
must be considered. In this regard, it is believed preferable that
potential sensing failures lead to inherent passivity of a defibrillating
device.
In many known VF detectors and automatic implantable defibrillators, the
primary detection schemes would result in active mode failures unless
other lock-out circuitry is provided. Examples are R-wave sensors,
pressure sensors, and elastomeric contraction sensors. In each case a
failure in the primary sensor would have the same inherent effect as
fibrillation, causing the automatic implantable device to fire, an active
failure.
There is accordingly a great need for a VF detector which is accurate in
its detection of VF or other life threatening arrhythmias, so that failure
modes may be passive. It is toward the development of a VF detector such
as this that the present invention is directed. The present invention is
directed more generally to the development of an accurate, simple detector
of cardiac arrhythmias which overcomes or eliminates the drawbacks of
known detectors.
SUMMARY OF THE INVENTION
The present invention relates to a system for measuring the electrical
activity of the heart, and which can reliably discriminate between
hemodynamically efficient and inefficient arrhythmias, being particularly
sensitive to ventricular fibrillation. Though presented as a part of an
automatic implantable defibrillator, it should be appreciated that the
present invention is not limited to this specific application. For
example, and as noted above, other arrhythmias or tachyarrhythmias can
easily be identified by utilizing the teachings of the present invention.
Customarily, the term electrocardiogram (ECG) implies the use of electrodes
on the body surface to obtain electrical signals indicative of heart
activity. The term electrogram, on the other hand, generally refers to
measurements made at the surface of the heart. As used herein, "ECG" is
defined broadly, and refers to any measurement of the electrical activity
of the heart, notwithstanding the source or technique of the measurement.
With the present invention, VF may be detected with a degree of accuracy
never before possible, and hence inherent passive failure modes can be
afforded. The inventive detector enjoys operation independent from the
concepts of QRS detection and heart rate calculations to maximize
accuracy. As is known, these concepts are particularly difficult to define
during ventricular fibrillation. Furthermore, the present invention
provides a technique for utilizing an ECG signal to derive substantially
more information about heart function than has heretofore been possible.
In its first and most basic embodiment, the inventive VF detector depends
for its operation upon the principle of probability density function.
Briefly, the probability density function defines the fraction of time, on
the average, that a given signal spends between two amplitude limits. It
has been noted that the probability density function of an ECG changes
markedly between ventricular fibrillation and normal cardiac rhythm.
Accordingly, VF can be detected by providing a mechanism for generating a
probability density function (or a portion thereof), or approximating one
or more points on the function. The entire probability density function
need not always be developed; rather, it is sometimes sufficient to
develop only particular values of the function at certain sampling points.
In its first embodiment, the present invention relates to a method and a
simple circuit for developing and utilizing an entire probability density
function curve, or for developing the function, and sampling the same only
at select points. In copending application Ser. No. 878,006, on the other
hand, there is disclosed a simple circuit for detecting VF through the
means of what can be thought of as approximating a probability density
function at a single point, the ECG base line.
The present invention also relates to a VF detector which senses the
regularity of the R-to-R interval. It has been observed that during high
rate tachyarrhythmias (on the order of 250 beats per minute), R-waves can
still be identified, and almost always occur at a stable rate. During
fibrillation, on the other hand, there are no such regular R-waves.
Accordingly, the present invention relates to a detector, preferably
serving as a second-stage detector utilized with the probability density
function detector discussed above, includes a phase lock loop circuit
which monitors the variability in the R-to-R interval. The loop locks onto
regularly occurring R-waves, but if the R-to-R interval becomes irregular,
as in VF, the loop cannot lock. The present invention makes use of this
inability to lock as evidence of fibrillation.
Further considered in the present invention is a VF detector in the form of
an impedance sensor which measures impedance between cardiac electrodes.
Preferably, this detector serves as a second-stage detector to verify a VF
diagnosis from the inventive probability density function detector. It has
been found that the impedance due to cardiac contractions is related to
stroke volume. During fibrillation, stroke volume goes substantially to
zero, and a severe drop in pulsatile impedance change can be detected as
indicative of VF.
The impedance sensor discussed above requires a relatively large input
power to perform its sensing function. The present invention therefore
also contemplates a circuit by which the impedance sensor remains in a
stand-by low power state for the greater majority of time, and is fully
enabled only upon its preceding detector stage sensing what is diagnosed
to be VF.
It is accordingly the main object of the present invention to provide an
accurate detector of cardiac activity.
Another object of the invention is to provide a method and apparatus for
monitoring heart activity.
An added object of the invention is to provide a method and apparatus for
monitoring heart activity, for detecting abnormalities, and for
cardioverting a malfunctioning heart, when appropriate.
A further object of the present invention is to provide a system wherein a
probability density function generated in dependence upon an ECG trace is
utilized to provide information about cardiac activity.
Another object of the present invention is to provide a system utilizing
the principle of probability density function for detecting ventricular
fibrillation.
Still a further object of the present invention is to provide a phase lock
loop detector for detecting cardiac arrhythmias.
Another object of the present invention is to provide a two-stage detector
including a probability density function detector and a phase lock loop
detector.
Another object of the present invention is to provide a system for
detecting cardiac arrhythmias including an impedance sensing detector.
A further object of the present invention is to provide a two-stage VF
detector using a probability density function detector and an impedance
detector.
Still another object of the present invention is to provide a multi-stage
system for detecting cardiac arrhythmias, wherein one stage is fully
enabled only after a preceding stage has diagnosed an arrhythmia.
Another general object of the present invention is to provide a simple and
yet reliable VF detector which, should it fail, fails in a passive mode.
These and other objects of the present invention, as well as many of the
attendant advantages thereof, will become more readily apparent when
reference is made to the following description, taken in conjunction with
the accompanying drawings.
BRIEF DESCRIPTION OF THE DRAWINGS
FIG. 1(a) is a tracing of a square wave given for exemplary purposes;
FIG. 1(b) is a plot of the probability density function of the wave
illustrated in FIG. 1(a);
FIG. 2(a) is a typical catheter-sensed ECG trace;
FIG. 2(b) is a plot of the probability density function of the ECG trace
illustrated in FIG. 2(a);
FIG. 3(a) is an ECG trace representing ventricular fibrillation;
FIG. 3(b) is a curve representing the probability density function of the
ECG trace illustrated in FIG. 3(a);
FIG. 4 is a block diagram of a circuit for developing probability density
function traces for the input of an oscilloscope and/or for providing a
sampling function;
FIG. 5 is a block diagram of the inventive phase lock loop detector, shown
as a second-stage detector used in conjunction with a probability density
function detector;
FIGS. 6(a) through 6(f) represent signals at select locations in the
circuit illustrated in FIG. 5;
FIG. 7 is a block diagram of the inventive impedance sensor for detecting
ventricular fibrillation, shown as a second-stage detector used in
conjunction with a probability density function detector; and
FIGS. 8(a) through 8(f) are traces for explaining the operation of the
impedance detector illustrated in FIG. 7.
DETAILED DESCRIPTION OF THE DRAWINGS
The probability density function cardiac arrhythmia detector forming a part
of the present invention will first be described. However, before
embarking upon a detailed explanation of the inventive circuit, there
follows a brief discussion of the theory of probability density.
The inventive detector system is based upon a series of measurements on the
ECG. The measurements are known in the literature as the probability
density function, denoted as K.sub.x (X). If X(t) is a function of time,
then K.sub.x (X) can be interpreted as a function that defines the
fraction of time, on the average, that X(t) spends between two limits. For
example, the area under K.sub.x (X) between X=X.sub.1 and X=X.sub.2 is the
fraction of time that X(t) spends between the limits X.sub.1 and X.sub.2.
Looking at the simplified example illustrated in FIG. 1(a), it can be seen
that X(t) is always either at the levels X=-B or X=A, and that the
waveform spends half of its time at each one of these limits. The
probability density function for this example is illustrated in FIG. 1(b),
wherein the continuous function of time X(t) has been mapped into a
function of the amplitude-time distribution of X(t).
In developing the present invention, it was recognized that the shape of 1
probability density function of an ECG changes markedly between normal
cardiac rhythm and ventricular fibrillation. In this regard, the attention
of the reader is directed to FIG. 2(a) which illustrates a typical ECG
trace, to FIG. 2(b) which shows the probability density function of the
ECG illustrated in FIG. 2(a), to FIG. 3(a) which illustrates an exemplary
ECG trace representing ventricular fibrillation, and to FIG. 3(b) which is
the probability density function of the trace illustrated in FIG. 3(a). It
will be noted that when comparing normal cardiac rhythm with ventricular
fibrillation, the greatest changes occur in the respective ECG traces at
X=0, or at the baseline of the ECG signal. This is markedly reflected in
the probability density functions as can be seen when contrasting FIGS.
2(b) and 3(b).
In the most simplified embodiment of the present invention, as is more
fully described in the aforementioned copending application Ser. No.
878,006, an approximation of the probability density is developed for one
value of X, namely X=0 or at the baseline of a filtered ECG. It should be
noted that many other sampling levels are available for X other than zero,
as will be explained below, and hence the number or level of sampling
points are not in any way intended to be limited.
The probability density function provides another tool for viewing the
original time-amplitude function. All of the discrete characteristics of
the original signal are retained, but are displayed in a different format.
Thus information of general diagnostic significance is inherent in the
presentation and in some instances can more readily be seen or measured
automatically. The attention of the reader is therefore directed to FIG.
4, which illustrates a circuit, in greatly simplified block form, which
can be used to provide complete displays of probability density functions.
These traces of probability density provide a great deal of information in
the detection and study of tachyarrhythmias.
In FIG. 4, an input signal is introduced at input terminal 62, to then be
passed through an automatic gain control circuit 64. In this manner, input
signals of different amplitude can be handled by the overall circuit. On
the probability density display, signal amplitude appears on the abscissa,
and therefore, AGC will normalize the width of the display.
A digital storage element 66 follows AGC circuit 64, and serves to provide
a repetitive source of the input signal. The storage element 66 stores
approximately two seconds of ECG data in a digital memory, and continually
repeats this data. In this manner, the same data is repeatedly provided to
a window comparator 68.
The window comparator 68 provides a logical "1" whenever its input signal
lies within a narrow band centered around a band center "X". Then, by
passing the output of the window comparator 68 through a simple low-pass
filter 70, a voltage is developed which is proportional to the average
time that the input stays within the designated band. This signal is fed
to the "vertical" input of oscilloscope 72. And this is precisely
analogous to the definition of the distribution as defined above. Slowly
sweeping the band center "X" through the range of the input signal by
means of a wave generator 74, provides a continuous display on the
oscilloscope screen. The band center is coupled into the "horizontal"
input of the oscilloscope 72.
The respective traces of FIGS. 2(b) and 3(b) were developed from the
circuit illustrated in FIG. 4. The trace of FIG. 2(a), as noted above,
represents an electrogram recorded from an intracardiac catheter. The
corresponding density function is shown in FIG. 2(b). Several regions have
been identified on the respective traces, representing the same cardiac
events but on the two different display patterns. For example, the region
"B" of FIG. 2(a) is the most negative peak of the R-wave. The signal
spends very little time at this value, thus the corresponding peak of the
probability density curve of FIG. 2(b) is small. The peak at "A" is
representative of the ST segment, and as is readily seen, the ECG signal
spends more time on the average at this level than at region "B".
Accordingly, the peak is higher at "A" in FIG. 2(b). The ECG dwells
longest at the baseline identified at "C" in FIG. 2(a), and the zero peak
is hence the largest in FIG. 2(b). FIGS. 2(a) and 2(b) also show an ECG
peak at "D", and the correspondence of this peak on the probability
density function curve. In accordance with the teachings of the present
invention, the absence of a peak at zero in the probability density
function is used as being characteristic of abnormal cardiac rhythm.
As mentioned previously, it is possible to detect VF by sampling points on
a probability density function, rather than by developing and monitoring
the entire function. For example, if two sampling points, say X.sub.1 and
X.sub.0, are defined as shown in FIG. 3(b), more discrimination resolution
becomes available by taking a ratio. As illustrated, approximate
measurement at these two points on the waveform for this example would
show the value of the probability density function to be:
For Normal Rhythm
##EQU1##
For Ventricular Fibrillation
##EQU2##
It can readily be seen that this measure yields over two orders of
magnitude difference between normal cardiac rhythm and fibrillation.
Sensing a value of C.sub.m near 1.0 corresponds to the detection of a
severe arrhythmia. Sampling at two portions of the probability density
function is indicated in FIG. 4 at 75.
With continuing reference to FIG. 4, it can be seen that the monitored
probability density function can be used to initiate the discharge of
energy into a needy heart. Although illustrated only in connection with
X.sub.0 and X.sub.1 sampling, cardioversion can be controlled through the
portion of the monitoring circuit which develops the complete probability
density function.
In FIG. 4, a known input signal, corresponding to a normal probability
density, is introduced to a comparator 50 at lead 52. The remaining input
54 to comparator 50 receives signals from X.sub.0, X.sub.1 sampler 75. If
the input from sampler 75 is normal, then comparator 50 remains idle. If,
on the other hand, a condition in need of cardioversion is indicated,
comparator 50 issues a "cardioversion indicated" signal, and discharge of
electrical energy into the heart is effected by discharge unit 56.
In view of the high degree of reliability necessary for the successful
application of an implantable automatic defibrillator, it may become
desirable to improve the accuracy of the detection system even relative to
that described immediately above. This can be done by adding stages of
sensing devices responsive to other parameters. One such parameter which
can aid in the discrimination of very severe tachyarrhythmias and
fibrillation is the variability in the R-to-R interval. As noted above,
even during extremely high-rate tachyarrhythmias, R-waves can be
identified and generally occur at a stable rate. During fibrillation, on
the other hand, all regularity in the R-waves is lost. It is therefore
possible to discriminate between fibrillation and tachyarrythmias by
monitoring the regularity of the R-wave intervals by means of an R-wave
detector.
In this latter regard, the present invention relates to the technique of
ascertaining R-to-R interval variability by way of a phase lock loop
circuit. The phase lock loop circuit has the capability of "locking" onto
periodic input signals and providing an AC output voltage which is at a
constant phase and at a frequency which is an integral multiple of the
input frequency. If the input is not periodic, however, the loop cannot
"lock", and this condition is easily detected. Accordingly, the inability
of the phase lock loop circuit to lock can be utilized to derive useful
information.
By utilizing the probability density function detector as a first detector
stage and a phase lock loop detector as a second detector stage, the
absence of a locked stage in the phase lock loop detector, coupled with
the condition of the first detector stage having issued a fibrillation
output, verifies the presence of VF with an exceedingly high degree of
accuracy. Phase lock loop circuits are well known in the literature, and
an example of a low-power version with lock indication, directly
applicable to fibrillation detection, can be found in Application Note
ICAN-6101, RCA COS/MOS Integrated Circuits, 1975 Databook Series, pp.
471-478. Accordingly, the phase lock loop circuitry employed in the
present invention is shown only in block form in FIG. 5. Its application
to a fibrillation detector, and the use of inability to lock to derive
information are, however, novel concepts.
With reference now to FIG. 5 and FIGS. 6(a)-6(f), the inventive use of a
phase lock loop circuit in a fibrillation detection system will be
described. Although the phase lock loop detector can alone be used to
detect VF, it is illustrated in FIG. 5 as a second-stage of a VF detector.
In FIG. 5, the previously-discussed probability density function
fibrillation detector is illustrated as an integral part of a first stage
detector shown at 76. The input to the first stage detector 76 reaches an
ECG amplifier 78, and is processed by the probability density function
detector 80. If fibrillation is sensed by the detector 80, then a signal
is issued at line 82, and is fed to one terminal of an AND gate 84. The
second input terminal 86 of AND gate 84 is associated with the second
stage of the inventive detector combination, and in particular, the phase
lock loop circuit shown generally at 92.
The signal issued by the ECG amplifier 78 also serves as an input to a
filter 88 which feeds filtered signals to an R-wave detector 90, each
being of conventional design. The ECG signal to filter 88 is illustrated
in FIG. 6(a), while the filtered signal serving as the input to the R-wave
detector 90 is shown in FIG. 6(b).
The R-wave detector 90 senses the presence of R-waves, and for each R-wave,
issues a pulse of finite period. If the R-waves are regular in interval,
the output of the R-wave detector 90 is a periodic train of pulses. FIG.
6(c) illustrates the output of the R-wave detector 90 based upon the input
of the filtered ECG shown in FIG. 6(b). It will be noted that the first
three pulses of R-wave detector 90 are periodic.
The phase lock loop circuit 92 includes a phase detector 94, the output of
which is filtered by a low-pass filter 96, in turn feeding signals to the
control side of a voltage controlled oscillator 98. The oscillator 98
issues a regular train of square wave pulses and feeds the same to the
phase detector 94, which then compares the phase of these regular pulses
with the input from the R-wave detector 90.
The phase lock loop 92 may be of numerous designs, as these circuits are
well known. In any event, it is contemplated in the present invention that
the phase detector 94 provide output information for a lock detector 100
which is indicative of the phase relationship between the R-wave detector
pulses and the oscillator pulses, and in turn, which indicates whether the
phase lock loop 92 is able to lock onto the input from the R-wave detector
90.
Upon the lock detector 100 receiving an indication from the phase detector
94 that the loop is locked, the detector 100 will, for example, issue a
logical "0". Under this set of conditions, the AND gate 84 will remain
idle, even if the probability density detector 80 has indicated on lead 82
that fibrillation is present. On the other hand, if the lock detector 100
receives a signal indicating that the loop 92 is not locked, then a
logical "1" will be issued to AND gate 84, and if the logical "1" occurs
simultaneously with a similar signal from the probability density detector
80, then the AND gate 84 will issue a signal on line 102 which will
trigger the defibrillating electronics and actuate an energy delivery
mechanism 103. Still, however, if the phase lock loop 92 cannot lock, and
yet the probability density detector 80 has sensed no irregularity, then
the AND gate 84 will remain idle. This is illustrated in FIGS. 6(d)
through 6(f). FIG. 6(d) shows the output of the probability density
function detector 80, FIG. 6(e) represents the output of the lock detector
100, and FIG. 6(f) represents the output of the AND gate 84 at lead 102.
As noted above, the phase lock loop circuit has before been used in cardiac
applications. Typically, the loop is designed to lock onto the R-wave
signals, even if such signals vary over wide ranges of input conditions.
See, for example, the Van Horn et al U.S. Pat. No. 3,811,428. There, the
phase lock loop circuit is used to lock onto fetal R-waves so that even
during such times when the fetal heart beat is masked by the maternal
heart beat, it is possible to provide an indication of predicted fetal
beat. With the present invention, for the first time, the inability of a
phase lock loop circuit to lock provides useful information; when the
circuit is unable to lock, VF is indicated, and appropriate action is then
taken.
Now, the inventive impedance sensor VF detection circuit will be described.
It has already been proposed to use intraventricular pressure as an input
variable to a fibrillation sensor. See U.S. Pat. Nos. Re. 27,757 and Re.
27,652, assigned to the present assignee. A drop in ventricular pressure
corresponds to a life threatening arrhythmia, and this measurement is
advantageous for it provides a direct indication of the life threatening
condition, namely a loss of cardiac output. Unfortunately, long-term
reliable pressure measurements are difficult to make. The inventive
impedance detector is able to provide the same direct measurement of the
life threatening arrhythmia, and is also capable of accurate long term
utilization.
An electrical measurement which is easily and reliably made is the
electrical impedance between two electrodes in or about the heart. It has
been found that this impedance changes in accordance with cardiac
contractions, and is directly related to stroke volume. That is, the
impedance between cardiac electrodes varies in accordance with the volume
of blood in the heart. When in normal rhythm, the heart regularly
contracts and fills, and hence the normal impedance change is periodic.
During fibrillation, however, stroke volume essentially goes to zero, and
hence a severe drop in pulsatile impedance change can be seen. The
inventive circuit illustrated in block form in FIG. 7 is able to detect
the absence of pulsatile impedance changes, analogous to a drop in stroke
volume, and hence ventricular pressure. The traces of FIGS. 8(a) through
8(f) relate to the circuit illustrated in FIG. 7.
With reference then to FIGS. 7 and 8, the inventive impedance VF detector
is shown generally at 104. The detector 104 is powered by a power supply
105 on line 106, placed in the circuit by a gate 108 which is, in turn,
controlled by the probability density function detector shown at 110. It
is possible to utilize the impedance detector alone to recognize and
respond to VF. However, as noted previously, the impedance VF detector
requires a substantial amount of power from the implanted battery source.
Therefore, so as not to drain the battery, the dector 104 preferably
serves as a second-stage detector, and is so illustrated in FIG. 7. When
used as a second-stage detector, the impedance detector 104 is designed to
remain in a stand-by state until the probability density function detector
110 senses an abnormaility and triggers the impedance detector 104 by
placing its power supply 105 into the impedance detector circuit 104
through the means of gate 108. In this way, the circuit of FIG. 7 provides
an implied "AND" function. That is, the second-stage circuit 104, which
triggers the defibrillating electronics, is only actuated upon command
from the first-stage probability density function detector 110. Therefore
both circuits must agree that fibrillation is present before a
fibrillation output is generated.
The basic element in the impedance VF detector 104 is illustrated
schematically as cardiac condition impedance 112. The impedance 112 is,
for example, related to the impedance of the blood and tissue measured
across intracardiac electrodes spaced apart on a catheter. A current
source 114 associates with the impedance 112 and provides a current input
of constant value. An oscillator 116 feeds the current source 114 so that
source 114 generates an AC current to the impedance 112. In this manner,
the voltage across the impedance 112 will be proportional to the current
multiplied by the impedance value. As typical values, the oscillator 116
is set to 100 KHz, with the current source 114 supplying 100 .mu.a. The
impedance 112 is typically on the order of 50 ohms, and therefore
approximately 5 mV appears across impedance 112. The voltage across the
impedance 112 is then amplified by means of a voltage amplifier 118, and
the amplified voltage from amplifier 118 is then demodulated by means of a
synchronous demodulator 120.
The amplified and demodulated output of demodulator 120 is fed to a
bandpass amplifier 122, and then to a trigger network 124, a ramp
generator 126, and a threshold detector 128. The output of the threshold
detector 128, if present, appears at terminal 130, and serves to trigger
the defibrillation circuitry into operation; this is indicated by the
energy delivery block 131.
FIG. 8(a) represents an ECG which is at first normal, and which then
indicates fibrillation; FIG. 8(b) shows, in an exaggerated form so as to
appear on the same time scale, the output of oscillator 116; and FIG. 8(c)
represents a trace of the voltage across impedance 112 after amplification
by amplifier 118 and corresponding to the ECG in FIG. 8(a). It can be seen
in FIG. 8(c) that the voltage across impedance 112 increases for each
normal beat of the heart as blood is ejected from the heart.
The output of demodulator 120, after amplification by amplifier 122, is
illustrated in FIG. 8(d) where a negative-going signal is indicated for
each reduction in voltage, or pulse, across impedance 112. Ramp generator
126 develops a ramp which is shown in FIG. 8(e). It will be noted that the
ramp returns to its baseline each time the demodulated and amplified
output of amplifier 122, represented in FIG. 8(c), crosses a set threshold
level. Accordingly, during normal cardiac rhythm, the threshold detector
128 remains inactive. However, once fibrillation commences, where
indicated in FIG. 8(a), the curve of FIG. 8(d) smoothes out, without the
threshold being reached, and therefore the ram | | |
